Cardiac arrhythmia detection system for an implantable stimulation device

ABSTRACT

A cardiac event and arrhythmia detection system and method detects arrhythmic cardiac activity or other information from an electrogram signal of a heart. The system senses the electrogram signal through an electrogram lead, preliminarily processes the signal, and converts it to a plurality of discrete digital signals, each of which represents the magnitude of the electrogram signal at a prescribed sample time. The discrete digital signals are applied to both a cardiac event detector which has a dynamic threshold which is programmably adjustable so that T-waves are not sensed and a morphology detector. The morphology detector detects selected changes in the morphology (shape) of the electrogram signal, wherein such changes automatically control the sensitivity (gain and/or threshold) used to detect cardiac events. The occurrence of a prescribed amount of change in the detected morphology over time indicates the occurrence of a prescribed arrhythmic cardiac condition.

CROSS-REFERENCE TO RELATED APPLICATIONS

This is a continuation-in-part of application Ser. No. 08/310,687, filedSep. 22, 1994 now U.S. Pat. No. 5,513,644, which is acontinuation-in-part of application Ser. No. 07/984,157, filed Dec. 1,1992, now abandoned.

FIELD OF THE INVENTION

The present invention relates to a cardiac arrhythmia detection systemfor use in an implantable stimulation device, such as an implantablepacemaker, cardioverter or a defibrillator. More particularly, thepresent invention relates to a system for detecting and tracking cardiacevents and electrogram morphology so that the transition betweenrhythmic activity to arrhythmic activity can be detected. The presentinvention further includes an improved implantable cardiac eventdetection system which eliminates double sensing of T-waves.

BACKGROUND OF THE INVENTION

The major pumping chambers in the human heart are the left and rightventricles. The simultaneous physical contraction of the myocardialtissue in these chambers expels blood into the aorta and the pulmonaryartery. Blood enters the ventricles from smaller antechambers called theleft and right atria which contract about 100 milliseconds (ms) beforethe ventricles. This interval is known as the atrioventricular (AV)delay. The physical contractions of the muscle tissue result from thedepolarization of such tissue, which depolarization is induced by a waveof spontaneous electrical excitation which begins in the right atrium,spreads to the left atrium and then enters the AV node which delays itspassage to the ventricles via the so-called bundle of His. The frequencyof the waves of excitation is normally regulated metabolically by thesinus node. The atrial rate is thus referred to as the sinus rate orsinus rhythm of the heart.

Electrical signals corresponding to the depolarization of the myocardialmuscle tissue appear in the patient's electrocardiogram. A brief lowamplitude signal known as the P-wave accompanies atrial depolarizationnormally followed by a much larger amplitude signal, known as the QRScomplex, with a predominant R-wave signifying ventriculardepolarization. Repolarization prior to the next contraction is markedby a broad waveform in the electrocardiogram known as the T-wave.

A typical implanted cardiac pacer (or pacemaker) operates by supplyingmissing stimulation pulses through an electrode on a pacing lead incontact with the atrial or ventricular muscle tissue. The electricalstimulus independently initiates depolarization of the myocardial(atrial or ventricular) tissue resulting in the desired contraction. TheP-wave or R-wave can be sensed through the same lead (i.e., the pacinglead) and used as a timing signal to synchronize or inhibit stimulationpulses in relation to spontaneous (natural or intrinsic) cardiacactivity. The sensed P-wave or R-wave signals are referred to as anatrial electrogram or ventricular electrogram, respectively.

Note that the term electrogram lead is used herein to refer to the leadthat transmits the sensed electrogram signal from the heart, and theterm pacing lead is used to refer to the lead that transmits thestimulation pulse to the heart. As mentioned above, however, these"leads" are generally combined (i.e., the sensed electrogram signal istransmitted from the heart by the same lead that transmits thestimulation pulse to the heart). The separate terms "electrogram lead"and "pacing lead" are used herein merely to indicate that theelectrogram signal and the stimulation pulse could be transmitted usingseparate leads.

Every modern-day implantable pacemaker includes a sensing circuit,whether the activity of one or both chambers of the heart are sensed. Acardiac event is sensed when an amplified electrogram signal exceeds athreshold value. If the sensitivity level is too low (i.e., the gain istoo low), then some cardiac events will not be sensed because even peaksignals may not exceed the threshold level. If the sensitivity level istoo high, on the other hand, the high gain of the amplifier may causenoise or T-wave signals to be sensed, giving rise to erroneous sensingof cardiac events. Pacemakers provided with communications telemetry(e.g.,noninvasive programming capabilities) advantageously allow thephysician to set the sensitivity level.

There are at least two disadvantages to having the physician set thesensitivity level. First, adjusting the sensitivity level is one morething that the physician must remember to do, and it would beadvantageous to relieve him or her of that task if it is possible to doso. Second, and more important, the physician generally sees the patientonly occasionally, and weeks or months may go by without the sensitivitylevel being changed. Problematically, the sensitivity level that willaccurately detect cardiac events at a given threshold level for apatient does not stay static; both R-wave amplitude and frequencycontent can vary considerably within a given patient. Changes in thesensitivity level are needed to accommodate for physical and mentalstress. In addition, the sensitivity level needs to change as myocardialtissue (heart muscle tissue) undergoes scarring or other physicalresponses to the implanted electrogram lead(s). other changes in themyocardium-electrogram lead interface, e.g., shifting of the position ofthe electrogram lead, may also cause changes in the proper sensitivitylevel.

Unfortunately, these changes occur over a period of days and, in somecases, even hours or minutes. Because the physician generally sees thepatient only every few weeks or months, the pacemaker sensing circuitscan erroneously detect, or not detect, cardiac events over large periodsof time. This erroneous detection/non-detection can cause under-pacingor over-pacing of the heart. Unfortunately for the patient, such changesmay potentially leave him or her in a worse condition than he or she wasin before the pacemaker was implanted. At best, the pacemaker is notable to operate efficiently--either by unnecessarily pacing and therebydraining the battery and risking pacemaker-induced tachycardias; or bynot pacing as often as is needed by the patient. Thus, what is needed isa way to adjust the sensitivity level of a cardiac event detector inresponse to changing conditions in an electrogram signal over a shortperiod of time.

One way of adjusting the sensitivity level of a cardiac event detectoris discussed in U.S. Pat. No. 4,708,144 issued to Hamilton et al. TheHamilton et al. patent shows the use of an attenuator to attenuate anamplified signal before such signal is digitized and rectified. Afterthe signal is digitized and rectified, the signal is connected to adigital comparator. The digital comparator compares each digitized andrectified sample to a threshold value. If the digitized and rectifiedsample exceeds the threshold value, a cardiac event is detected. Inresponse to the detected cardiac event, a pacemaker control circuittakes appropriate action.

Each digitized and rectified sample is also presented to a peak detectorwhich stores the maximum, or largest, digitized and rectified sample.The stored maximum sample is coupled to the pacemaker control circuit.After each cardiac event is sensed, the pacemaker control circuitaverages the stored maximum with any of the previously occurringmaximums yielding an average peak value. This average peak value is usedto determine whether or not the attenuator should be adjusted toincrease or decrease the attenuation provided by the attenuator.Specifically, if the average peak value increases, the attenuation isincreased; and if the average peak value decreases, the attenuation isdecreased, thereby adjusting the amplitude of the signal before it isdigitized and rectified.

Disadvantageously, even though the Hamilton et al. circuit provides onetechnique for dynamically adjusting the sensitivity level of a cardiacevent detector, by attenuating the cardiac signals after theamplification stage, it suffers from potentially clipping the inputsignal before even reaching the attenuation stage or the peak detectingstage. Furthermore, it completely lacks the ability to eliminate thesensing of high amplitude T-waves, which could cause the peak detectorto erroneously detect the peak T-wave. Furthermore, if the amplitude ofthe T-wave is too high, or if the gain of the amplifier is so high thatthe R-wave is clipped and the T-wave is, by comparison, similar inamplitude to the clipped R-wave, it can result in "double sensing".Double sensing, when it occurs, then falsely indicates that atachycardia is present. Thus, Hamilton et al. is not suitable forimportant cardiac monitoring functions beyond merely sensing a cardiacevent. What is needed is a system which adjusts the gain at thepre-amplification stage for an optimum signal (i.e., without clippingthe input signal) and then reliably eliminate sensing of the T-wave.

In addition to the detection of cardiac events, it is desirable, in thetreatment of certain heart ailments, or for the detection of suchailments, to continuously monitor the patient over a certain period oftime in order to determine the effectiveness of the treatment beingadministered by a pacemaker, under different conditions of stress orvarying conditions of the heart. If the sensing circuit detects that thepacemaker is administering a less than ideal, or optimum, treatment, thetreatment can be adjusted (e.g., by increasing or decreasing the rate atwhich pacing pulses are delivered, by decreasing the threshold level ofthe threshold detector, or by increasing the amplitude or duration ofthe pacing pulse).

Unfortunately, events that would indicate that the pacemaker may beproviding less than the optimum treatment may occur only infrequently.Thus, a physician may not detect such abnormal events during a weekly,biweekly or monthly examination, which may last only a few minutes andmay not be able to adjust the pacemaker accordingly. In an effort tosolve this problem, data acquisition systems have been developed thatrecord electrogram signals over a predetermined period of time, e.g., onthe order of days. The electrogram signals may then be analyzed by aphysician or, in more advanced system, by a microcontroller in thepacemaker in accordance with a control program that is designed to reactto various conditions that are manifested by the electrogram signals.Such data acquisition systems advantageously allow detailed analysis ofthe electrogram signal over long periods of time thereby facilitatingthe detection and accommodation of infrequent heart abnormalities or theearly detection of slowly developing heart ailments. Such long-termmonitoring, particularly where implemented in advanced programmedsystems, makes possible the purposeful and possibly automatic treatmentheart abnormalities long before the actual failure of a pacemaker toproperly service the heart. In addition, with such automated systems,therapies, such as antitachycardia pacing and defibrillation, can beperformed on the heart by pacing systems or dedicated defibrillatorsthat would otherwise not be able to be performed as quickly orautomatically.

Unfortunately, implanted data acquisition systems have heretofore onlybeen operable over a limited sample of the electrogram signal. This isbecause such systems store the electrogram signal in a memory. Thememory is of a limited size, and when the memory is full, either part ofthe previously recorded electrogram signal must be discarded to makeroom for new electrogram signal to be recorded, or the data acquisitionsystem must stop recording. In an effort to solve this problem, varioushigh capacity means of storing electrogram signals have been developedsuch as magnetic tape recording systems. For example, U.S. Pat. No.4,250,888 issued to Grosskopf, suggests that when the memory is full, awarning message be given that alerts the patient to the need to contactthe physician or to activate a tape recording system at home.

Disadvantageously, the Grosskopf approach may require that the patientreport to a potentially inconvenient location, (i.e., the physician'soffice or the patient's home where the tape recording system islocated). Such inconvenience may encourage the patient to ignore thewarning message. In addition, the warning message can be intrusive andembarrassing. Furthermore, such warning systems are not used withimplantable pacemakers for at least two reasons. First, implantablepacemakers are implanted within the body and, as such, any warning meansare neither visible nor readily heard. Second, implantable pacemakersmust be compact and use little power. Generally, the warning message isgenerated by a speaker or light source and thus draws a significantcurrent. It is thus apparent that what is needed is an implantablecardiac event detection system that is not limited to operating on asmall sample of the cardiac signal over a limited period of time, andthat does not require the use of inconvenient and impractical storagedevices such as tape recording systems.

Some systems have been developed that store only anomalous portions ofthe electrogram signal. See, e.g., Grosskopf. However, even thesesystems have a limited capacity and when a sufficient number ofanomalous portions are stored, some data loss occurs. This data lossoccurs when the memory is full and either the new signal must bediscarded or the previously stored signal must be discarded.Problematically, the portion of the electrogram signal that is discardedmay be the portion of signal that is needed for an accurate evaluationof the patient's heart condition. Thus, what is needed is an implantablecardiac event detection system that is not limited by the use of afinite capacity memory for storing the electrogram signal, but thatprovides information sufficient for programmed evaluation in amicrocontroller and, if needed, automatic adjustment of a pacemaker oractivation of a defibrillator in response to such evaluation.

Another problem faced by the designers of automated cardiac pacingand/or defibrillation systems is the need for analysis of theelectrogram signal. One approach to accurately analyzing the electrogramsignal requires that the stored electrogram signal be subjected tocomplex digital filtering algorithms and statistical analysis. See e.g.,U.S. Pat. No. 4,422,459 issued to Simson. In order to generate thedigitally filtered and statistically analyzed signals in Simson, a largecomputer system is employed. Such computer system is immobile andinconveniently located at, e.g., the physician's office, thus makingimplantation impossible. Disadvantageously, such algorithms and analysisrequire that many hundreds of mathematical operations be performedbefore an accurate conclusion as to whether the cardiac pacer and/ordefibrillator are performing optimally can be obtained and, thus, beforeneeded adjustment of the therapies provided by the cardiac pacer and/ordefibrillator can be made. Problematically, this requires not only theuse of a memory to store the incoming electrogram signal while themathematical operations are being completed, the disadvantages of whichare discussed above, but requires that many complicated computationalsteps be traversed by the microcontroller. Such complicatedcomputational steps are highly power-consuming--which would require morefrequent replacement of the battery that powers the implantable cardiacpacer and/or defibrillator--and thus, disadvantageous in implantablecardiac pacing applications.

Another approach to accurately analyzing the electrogram signal has beento allow the physician to analyze the electrogram signal stored in amemory using conventional electrogram analysis techniques.Disadvantageously, in order to obtain the stored electrogram signal, thephysician must download the stored electrogram signal via a telemetrycircuit in the cardiac pacer system and/or defibrillator system. Hence,because a memory is used, the problems discussed above are also presentin this approach. A further disadvantage of this approach is that noautomated adjustment of the therapies provided by the cardiac pacerand/or defibrillator can be made because such adjustment must wait untilthe patient has traveled to the physician's office and until thephysician has completed his or her analysis. Thus, what is needed is animplantable cardiac data acquisition and analysis system that does notrequire the use of complicated and highly power-consuming mathematicalcomputations.

SUMMARY OF THE INVENTION

The present invention advantageously addresses the above and other needsby providing an improved implantable cardiac event detection system andmethod usable with implantable cardiac pacemakers, cardioverters,defibrillators, or the like.

One aspect of the present invention provides an implantable cardiacevent detection system which can eliminate the sensing of T-waves,reliably and dynamically set the detection threshold, and optimize thedynamic range of the incoming cardiac signal (through automatic gaincontrol of a pre-amplifier stage) for the purpose of automaticallyadapting to changing cardiac signal morphology.

The system is coupled to the heart via an electrogram lead as is knownin the art of cardiac pacing. The system is also coupled to a therapycircuit that provides pacing and/or defibrillation therapies to theheart. An electrogram signal is sensed through the electrogram lead andis transmitted to signal conditioning circuitry. The signal conditioningcircuitry includes a pre-amplifier having a plurality of programmablegains, a narrow band filter, and digitizing circuitry. The resultantdigitized electrogram signal is then coupled to a threshold detector andto a morphology detector, which are, in turn, coupled to amicrocontroller. The microcontroller controls the threshold value usedby the threshold detector, as well as controlling the gain of thepre-amplifier, as a function of the various morphology parameters (e.g.,the previous average peak R-wave value) sensed by the morphologydetector.

The threshold detector detects a cardiac event (e.g., an R-wave) withinthe electrogram signal whenever the electrogram signal exceeds aprescribed initial threshold value. In the preferred embodiment, thethreshold detector is a digital threshold detector capable of digitallyadjusting the threshold value in a predetermined stepwise fashion. Thethreshold detector eliminates the detection of high amplitude T-waves byincreasing the initial threshold value to a temporary threshold valuefor a prescribed, or programmable, period of time following thedetection of an R-wave. In an alternate embodiment, the temporarythreshold value is initiated after a prescribed, or programmable, delayperiod following the detection of an R-wave.

The threshold value is then gradually ramped down, in a stepwisefashion, to its initial value within a second prescribed period of time.Advantageously, the second prescribed time period may be automaticallyadjusted as a function of heart rate (e.g., a fast heart rate wouldrequire a fast ramp down to the initial value). In this manner, thedetection of T-waves are eliminated. Thus, "double sensing" of theT-wave and false indications that a tachycardia is present cannot occur.

In the preferred embodiment, the initial threshold value isautomatically determined by the microcontroller to be a percentage ofthe average peak (or maximum) R-wave signals over at least a period of afew minutes (e.g., at 25% of the previous peak values). Preferably, thetemporary threshold value is also automatically set to a percentage ofthe average peak (or maximum) R-wave signals over at least a period of afew minutes (e.g., at 100% of the previous average peak values).Alternately, the initial and the temporary values could be aprogrammable value.

In the preferred embodiment, the gain of the pre-amplifier isautomatically adjusted for the optimum dynamic range so that theincoming cardiac signal is not clipped. The control of the pre-amplifieris determined by the microcontroller and the morphology detector, asdescribed in more detail below.

Another aspect of the present invention provides an implantable cardiacarrhythmia detection system for detecting the transition betweenrhythmic and arrhythmic cardiac activity in a heart using the morphologydetector. One feature of the present invention is the detection of ashift in the average baseline of the rectified cardiac signal. Forexample, during normal sinus rhythm, the average baseline of the cardiacelectrogram is approximately zero. When an arrhythmia occurs (such as, atachycardia or fibrillation), the average baseline of the cardiacelectrogram increases in magnitude. It is this detection of the shiftingof the average baseline which is used, in the present invention, todetect a change in the patient's cardiac rhythm.

In one embodiment the "average baseline" may be thought of as an RMSvalue of the electrogram signal, or the average of the rectified signal,in that, only the positive values are considered. In the preferredembodiment, the "average baseline" is the sum of the unsigned magnitudesof a plurality of digitized samples during a prescribed interval.

Another feature of the present invention is the detection of a change inthe morphology of the R-wave as a way to indicate a change betweenrhythmic and arrhythmic cardiac activity. For example, if the amplitudeof the R-wave increase or decreases by a prescribed amount, or changespolarity, chances are that a new ectopic foci is generating R-waves froma new location, thereby indicating a change in the patient's cardiacrhythm. Furthermore, when using morphology changes in combination withthe shifting of the average baseline, an even higher confidence level isachieved that the patient's cardiac rhythm has become arrhythmic.

Thus, in the present invention, the morphology detector detects variousparameters associated with the morphology of the electrogram signal, andcouples such parameters to the microcontroller. (Note, as used hereinthe term "morphology" relates to the shape of the electrogram signalwhen viewed as a signal waveform as a function of time.) Themicrocontroller then indicates the presence of an arrhythmic cardiaccondition in response to a prescribed change in the morphologyparameters. Such indication may then be used by the desired therapycircuit, e.g., a pacemaker, cardioverter or defibrillator, in order todeliver an appropriate therapy to the heart.

The morphology detector, according to the present invention, includesone or more of the following: a minimum detector, a maximum detector, apeak detector, a baseline averager, a baseline sampler, an accumulator,and/or an interval counter; the output signals of each being coupled tothe microcontroller for determining the presence of an arrhythmia asdescribed in more detail below.

The minimum detector, used in the morphology detector, generates aminimum signal, indicative of the magnitude of the most negative value(i.e., below a baseline voltage) in the electrogram signal and recordsthe magnitude of such value as a minimum value.

The maximum detector generates a maximum signal, indicative of themagnitude of the most positive value (i.e., above the baseline voltage)in the electrogram signal and records the magnitude of such value as amaximum value.

The peak detector, used in the morphology detector, generates a peaksignal, indicative of the largest value in the electrogram signal,regardless of whether the largest value is negative or positive.

Alternatively, the peak signal may be generated by the microcontroller,in which the microcontroller determines the peak signal to be thegreater of the minimum value and the maximum value. (Note that theminimum and maximum values once determined, are both considered aspositive magnitudes).

The baseline averager generates an "average baseline signal" indicativeof the average magnitude of the electrogram signal over a predeterminedtime period. The predetermined time period may be the time during whichthe baseline is expected to be quiescent (e.g., as determined by thecounter or by the microcontroller). Alternatively, the predeterminedtime period may be the entire cardiac cycle. The rational for the latteris that during normal sinus rhythm, the average baseline signal over theentire cardiac cycle approximates the true (quiescent) baseline.

As an alternative to the baseline averager, in accordance with oneembodiment of the invention, a baseline sampler may be used by themorphology detector. The baseline sampler generates an accumulatedbaseline signal, indicative of the accumulated magnitude (i.e., the sumof all baseline values) of the electrogram signal during thepredetermined period of time. The microcontroller then counts the numberof discrete sample values in the processed electrogram signal that areaccumulated by the baseline sampler during the predetermined timeperiod. In this way, a count value is generated. The total baselinevalue, from the baseline sampler, is then divided by such count value inorder to generate the average baseline value.

Typically, the baseline sampler operates digitally. That is, theelectrogram signal is sampled and digitized. The microcontroller thensimply reads the magnitude of the digitized samples occurring during thepredetermined time period and divides it by the total number of samples,as determined by the interval counter.

In accordance with another embodiment of the invention, the electrogramsignal is not digitized (i.e. the processed electrogram signal isanalog). The baseline sampler may analogically and rapidly sample themagnitudes of the processed electrogram signal over the predeterminedtime period. The magnitude of each of the rapid samplings (discretevalues) occurring during the predetermined time period are addedtogether. In this way, the accumulated baseline signal can beanalogically generated. The accumulated baseline signal is coupled tothe microcontroller, which then divides it by the time interval, asdetermined by, e.g., an interval counter.

The accumulator, when used within the morphology detector, generates anaccumulated magnitude signal indicative of the accumulated magnitude ofthe electrogram signal during a cardiac cycle. The cardiac cycle beginswhen an R-wave is detected by the threshold detector, as describedabove, and ends when a subsequent R-wave is detected. The accumulatormay operate digitally or analogically in the same or similar manner asthe baseline sampler described above. The accumulated magnitude signalmay then be divided by the total number of samples over the entirecardiac cycle to produce an average baseline signal over the entirecardiac cycle.

In one embodiment, the microcontroller provides a count signalcorresponding to the number of samples in the predetermined time period(or cardiac cycle). Alternatively, the count signal is determined usinga counter (as opposed to counting with the microcontroller). Such countsignal is then coupled to the microcontroller, where the accumulatedbaseline value from the baseline sampler is divided by such count valueto generate the average baseline value.

An arrhythmia flag signal is generated by the microcontroller when thequotient of the peak (or max) R-wave value divided by the averagebaseline value exceeds an arrhythmia threshold value (stored, e.g., inthe microcontroller). The arrhythmia flag signal may then be used toengage a therapy circuit, e.g., a circuit that issues stimulation and/ordefibrillation pulses in a prescribed pattern.

In addition, the microcontroller may generate one or more othermorphology change signals, such as signals indicating: a change inpolarity, the amplitude of the R-waves or the gain of the pre-amplifier(which is related to the envelope of the incoming cardiac signal). Acontrol program executed by the microcontroller controls the generationof such output signals. The control program is executed in response tothe detection of an R-wave, and/or a time-out signal (i.e., the time-outsignal is generated in the absence of cardiac signals in order to limitthe number of samples that will be acquired by the event detector).

In the discussion below, it is generally assumed that the maximum valueof the cardiac signal is greater than the minimum value; however, it isto be understood that in some instances the minimum value may be largerthan the maximum value. Thus, while the following discussion is directedto setting various morphology change values based on changes in themaximum value, in the event that the minimum value is greater than themaximum value, the morphology change values would instead be set inresponse to changes in the minimum value.

During normal sinus cardiac rhythm, the relationship between themagnitude of the maximum value and the magnitude of the minimum valuedoes not typically change (i.e., the maximum value generally remainslarger than the minimum value). In the event that the magnitude of theminimum value suddenly becomes larger than the magnitude of the maximumvalue, a radical change in the morphology of the electrogram signal isindicated. The microcontroller sets the morphology change value to"POLARITY" in response to such radical change. Furthermore, because itis likely that the signals thereafter generated by the morphologydetector and/or the microcontroller are no longer indicative of themorphology of the previous electrogram signal, the microcontrollerresets the morphology detector and/or the output signals (e.g., theminimum value, the maximum value, the peak value, and the baselinevalue, etc.) generated by the microcontroller in response to a change inthe morphology change value, e.g., when the polarity changes.

In addition to setting the morphology change value to "POLARITY," themicrocontroller may set the morphology change value to "INCREASE,""DECREASE," and/or "NONE." For example, in the event that the presentmaximum value (or minimum value) is much greater than the average of thepreceding maximum values (or minimum values), the microcontroller setsthe morphology change value to "INCREASE." When the morphology changevalue is set to "INCREASE," there is a high probability that a cardiacarrhythmia has begun. The therapy circuit may then respond to suchdetected arrhythmia as is known in the art of implantable pacemakers.

In the event that the present maximum value (or minimum value) is muchlower than the average of the preceding maximum values (or minimumvalues), but is still greater than the minimum value (or maximum value),the microcontroller sets the morphology change value to "DECREASE." Whenthe morphology change value is set to "DECREASE," there is a likewisehigh probability that a cardiac arrhythmia has begun, and the therapycircuit may thus respond accordingly.

In the event that the present maximum value (or minimum value) is notmuch greater or much less than the average of the preceding maximumvalues (or minimum values), and the present maximum value (or minimumvalue) remains greater than the present minimum value (or maximumvalue), the morphology change value is set to "NONE." This indicates tothe therapy circuit that the heart is experiencing normal sinus rhythm.In this way, changes in the morphology of the electrogram signal,particularly the onset of arrhythmias, can be detected without the needfor complicated and power-consuming computation by the microcontroller.

Note that in the event the peak detector is used instead of the minimumand maximum detectors, the morphology change value is set in response tochanges in the peak value in a manner similar to that described above.However, it is important to note that the morphology change value isgenerally not set to "POLARITY" when the peak detector is used insteadof the minimum and maximum detectors.

The gain change value output signal is generated by the microcontrollerin response to variations in an average of the previous maximum values(or minimum values). That is, in the event that the average of theprevious maximum values (or minimum values) increases substantially(e.g., as such average is updated over a period of time), themicrocontroller sets the gain change value to "DECREASE" and decreasesthe gain of the pre-amplifier (or other sense amplifier) used as a partof the signal conditioning circuitry. Similarly, if the average of theprevious maximum values (or minimum values) decreases substantially, themicrocontroller sets the gain change value to "INCREASE" and increasesthe gain of the pre-amplifier. If the average of the previous maximumvalues (or minimum values) does not change substantially, the gainchange value is set to "NONE" and the gain of the pre-amplifier is notchanged.

The gain change value and the control of the gain of the pre-amplifierserve two functions. First, a change in the gain change value isindicative of a change in the envelope of the cardiac signal, and thus,a change in the morphology. That is, a change in gain indicates thatR-waves are now being generated from a new ectopic foci. Therefore, thegain change value may be used to indicate an arrhythmia is present.Secondly, control of the gain of the amplifier helps to minimize thesensing of T-waves. That is, by maximizing the R-wave signal (bypreventing clipping), the effect is to minimize the T-wave amplitude,and further enables a larger range of thresholds which can detect theR-wave without detecting the T-wave. In this way, the gain of thepre-amplifier may be adjusted in response to varying amplitudes of theelectrogram signal.

In accordance with one aspect of the invention, the microcontroller setsthe sensitivity of the pre-amplifier based on the gain change value. Inthis way, the sensitivity of the pre-amplifiers can be set to an optimumdynamic range (i.e., without clipping the cardiac signal) so thatcardiac events are accurately sensed, while noise and the like are notmistaken for cardiac events.

The output signals of the microcontroller representing the averagebaseline value, the peak value, the arrhythmia flag value (i.e., thedetected shift in the average baseline signal or the peak value dividedby the average baseline), the morphology change value, and the gainchange value are used to generate the average baseline signal, anaverage peak signal, an arrhythmia flag signal, a morphology changesignal, and a gain change signal, respectively. Such output signals arecoupled to the appropriate therapy circuit, and are used by theimplanted therapy circuit to adjust the therapy services provided to theheart as needed.

It is thus apparent that important information is acquired from theelectrogram signal without the need for external storage devices, e.g.,magnetic tape recorders, and without the need for finite capacitymemories. It is also thus apparent that such information is provided tothe implanted therapy circuit without the need for complicated andhighly power-consuming mathematical calculations in an implantedmicrocontroller.

In addition to the threshold detector and to the morphology detector, atime-out counter may be used with the microcontroller. During normalsinus cardiac rhythm, the time-out counter receives a reset signal whenthe cardiac frequency has a period shorter than a predetermined periodof time. Therefore, the time-out signal is normally not generated.However, if the cardiac rhythm deviates from normal such that thecardiac frequency has a period longer than the predetermined period oftime, even for one period, (such as during an episode of bradycardia,asystole, sick sinus syndrome, etc.), then the time-out signal isgenerated. The microcontroller then executes the control program inresponse to the time-out signal (e.g., to generate the peak signal,arrhythmia flag signal, morphology change signal, etc.). In this way thecontrol program is executed when the cardiac frequency decreases below apredetermined frequency for even one period (in addition to beingexecuted when the event sensed signal is generated as described above).

Another aspect of the time-out counter is the ability to generate acycle length signal indicative of the period of the cardiac frequency.By dividing an accumulated magnitude value of the entire cardiac cycleby the cycle length signal, the microcontroller generates an averagemagnitude signal over the entire cardiac cycle.

A further aspect of the present invention provides a method of detectingthe transition between rhythmic cardiac activity and arrhythmic cardiacactivity. Such method includes the steps of: (a) sensing an electrogramsignal; (b) determining an average baseline value; and (c) detecting achange in the average baseline value by a prescribed amount, a changeless than the prescribed amount indicating rhythmic activity and achange greater than the prescribed amount indicating arrhythmicactivity.

In an alternative embodiment, the method includes the steps of: (a)sensing an electrogram signal; (b) recording at least one morphologychange value, (e.g., a change in polarity, a change in amplitude, achange in average baseline, or a change in gain); and (d) generating atleast one output signal indicative of any change in the morphologyvalue; which at least one output signal provides information about theelectrogram signal and may be used to control a therapy circuit.

It is thus a feature of the present invention to automatically adjustthe sensitivity of an implantable cardiac event detection system inresponse to varying electrogram signals without falsely detectingT-waves.

It is an additional feature of the present invention to easily andinexpensively acquire information from an electrogram signal over apotentially long (e.g., infinite) period of time.

It is a further feature of the present invention to acquire informationfrom an electrogram signal without having to use external storagedevices, e.g., magnetic tape recorders.

It is another feature of the invention to provide an implanted systemfor acquiring information from an electrogram signal without the needfor large memories to store all or portions of the electrogram signal.

It is yet a further feature of the invention to provide such animplantable information acquiring system that also generates outputsignals that may be used by an implantable therapy circuit, such as apacemaker, a defibrillator, or the like, and that facilitates automaticcontrol and adjustment of the therapy provided by such therapy circuit.

It is still an additional feature of the invention to provide such asystem wherein the output signals used to control the therapy circuitmay be generated without the need for performing complicated and highlypower-consuming mathematical computations in an implantedmicrocontroller.

BRIEF DESCRIPTION OF THE DRAWINGS

The above and other features of the present invention will become moreapparent from the Detailed Description of the Invention, presented inconjunction with the following drawings, wherein:

FIGS. 1, 2 and 3 show an intracardiac electrogram signal in normal sinusrhythm, tachycardia and in fibrillation, respectively, and illustratethe rational for using a shift in the average baseline to detect thetransition from rhythmic cardiac activity to arrhythmic cardiacactivity;

FIG. 4 is a block diagram showing a cardiac event and arrhythmiadetector of the present invention in conjunction with an electronictherapy circuit;

FIG. 5 is a basic block diagram of one embodiment of a cardiac event andarrhythmia detector made in accordance with the present invention;

FIG. 6 is a detailed block diagram of the cardiac event detector shownin FIG. 5;

FIG. 7 is a graph showing an exemplary electrogram signal used as aninput to the pre-amplifier of FIG. 6, and a representation of thedynamic threshold adjustment by the microcontroller;

FIG. 8 is a graph showing an exemplary electrogram signal used as aninput to the cardiac threshold detector of FIG. 5, and a digitizedelectrogram signal generated in response to the electrogram signal by asignal processor that is part of the morphology detector of FIG. 6; and

FIGS. 9, 10, 11, 12 and 13 are a flowchart of an exemplary controlprogram that may be used by a microcontroller that is part of thecardiac threshold detector of FIG. 6.

FIG. 14 is a graph showing an alternate embodiment of an electrogramsignal used as an input to the preamplifier of FIG. 6, and arepresentation of the dynamic threshold adjustment by themicrocontroller;

Like reference numerals are used to represent like elements in thevarious figures and the accompanying description that follows.

DETAILED DESCRIPTION OF THE INVENTION

The following description is of the best mode presently contemplated forcarrying out the invention. This description is not to be taken in alimiting sense, but is made merely for the purpose of describing thegeneral principles of the invention. The scope of the invention shouldbe determined with reference to the claims.

The basic premise for one aspect of the arrhythmia detection inventionof the present invention is illustrated in FIGS. 1-3. FIG. 1 illustratesa cardiac electrogram which is in normal sinus rhythm (NSR). FIG. 2illustrates a cardiac electrogram which, following a prematureventricular contraction (PVC), changes from normal sinus rhythm toventricular tachycardia (VT). FIG. 3 illustrates a cardiac electrogramwhich, following a premature ventricular contraction (PVC), changes fromnormal sinus rhythm to ventricular fibrillation.

The present invention detects the transition from rhythmic activity toarrhythmic activity by detecting a change in the baseline average. Inone embodiment, the baseline average is determined during a quiescentperiod of the signal, for example, during the time interval T₂,following a delay time interval, T₁, as shown in FIGS. 1, 2 and 3. Thedelay period, T₁, is chosen to begin after the repolarization signal(T-wave) has ended. In another embodiment, the baseline average isdetermined over the entire cardiac cycle (i.e., during the time intervalT₃), since this average closely approximates the true, or quiescent,baseline.

In the simplest embodiment, an increase in the average baseline beyond apredetermined threshold voltage will indicate a change in morphology,such as ventricular tachycardia and ventricular fibrillation, as shownin FIGS. 2 and 3, respectively. The predetermined threshold voltagewould be characterized at implant for a given patient.

In the preferred embodiment, to determine the shift in the baseline, theaverage peak (or max) R-wave value is divided by the average baselineand then compared to an arrhythmia threshold. If the result is greaterthan an arrhythmia threshold (also characterized at implantation), anarrhythmia flag signal is generated. For example, in FIG. 2 (or FIG. 3),after the delay period, T1, the system will begin to sample the baselineaverage during the time interval T2. At the end of T2, a higher baselinevalue is found as a result of the ventricular tachycardia (orventricular fibrillation). This higher baseline value is sent to amicrocontroller, which in turn, computes the quotient of the averagepeak R-wave value with the new baseline average. If the result isgreater than the patient's arrhythmia threshold, then the systemindicates that there is a high probability that an arrhythmia has begun.Furthermore, it may be possible to use the quotient of the average peakdivided by the baseline to discriminate between VT and VF.

Since the arrhythmia detection system of the present invention is basedon accurately detecting the peak amplitude of an R-wave, it can beappreciated by one skilled in the art that it is critical that T-wavesare not detected. Thus, an implantable cardiac event detection thateliminates T-waves is also described below.

In FIG. 4, a block diagram is shown of a cardiac detection system 21used with an electronic therapy circuit 23. The cardiac detection system21 and the therapy circuit 23 comprise, in combination, an implantabledevice 25 that is implanted into a patient and attached to a heart 10 ofthe patient via an electrogram lead 12. The electrogram lead provideselectrical communication between the implantable device 25 and the heart10. One signal that is communicated through the electrogram lead 12 fromthe heart 10 to the implantable device 25 is an electrogram signal. Suchelectrogram signals (sometimes referred to as EGM signals) are known inthe art. See, e.g., U.S. Pat. Nos. 4,596,255; 4,712,555; 4,817,605; and4,940,052; incorporated herein by reference. The electrogram signal iscoupled, within the implantable device 25, to the cardiac detectionsystem 21. The implantable device 25 is housed in an implantable,hermetically sealed housing as is known in the art of implantableelectronic pacemakers.

The cardiac detection system 21 processes the electrogram signal andgenerates at least one output signal. The output signal is coupled tothe therapy circuit 23, which therapy circuit may be considered as acardiac pacing system. The therapy circuit 23 controls the therapydelivered to the heart 10 (typically stimulation pulses) via theelectrogram lead 12. Alternatively, another lead (e.g., a pacing lead)may be used in lieu of or in addition to the lead 12 in order to deliverthe desired therapy (e.g., stimulation pulses) to the patient's heart.

By way of example, in FIG. 4, the therapy circuit 23 comprises a cardiacpacer. The therapy circuit 23 includes a control circuit 20, pulsegenerator circuitry 22, a memory circuit 24, and a telemetry circuit 26.In practice, the control circuit 20 (which may hereafter be referred tosimply as a "microcontroller") and the memory circuit 24 (hereafterreferred to as a "memory") may advantageously be utilized by both thecardiac detection system 21 and the therapy circuit 23, thereby assumingboth therapy-providing and event-detecting functions. However, to moreclearly illustrate the present invention, the control circuit 20 and thememory circuit 24 are shown as separate elements of the therapy circuit23, even though shared by the cardiac detection system 21. In someembodiments, the cardiac detection system 21 may utilize a separatemicrocontroller from the control circuit 20 and memory circuit 24 of thetherapy circuit 23.

The control circuit 20 receives the output signal from the cardiacdetection system 21 and, in response thereto, evaluates whether or notoptimum therapy is being delivered to the heart 10. If the therapy beingdelivered is not optimum, the control circuit 20 makes adjustments, asrequired. In response to the adjustment of the therapy, the pulsegenerator circuitry 22 begins delivering a more optimum therapy to theheart 10. Several adjustments of the pulse generator circuit 22 may beneeded before an optimum therapy is delivered. Under some circumstances,less than optimum therapy may be the best therapy that can be delivered.This is because the output signal may rapidly change in response tochanging conditions of the heart 10 before needed adjustments can bemade in response to previous output signals. However, the therapydelivered to the heart 10 will, in general, be repeatedly adjusted untiloptimum or near optimum therapy is delivered to the heart 10.

In determining what is optimum therapy, based on the output signalgenerated by the cardiac detection system 21, the control circuit 20 mayalso use a memory circuit 24. Various control parameters are stored inthe memory circuit 24 by a physician using a telemetry circuit 26. Inorder to store such parameters, the physician utilizes an external(non-implanted) programmer 28 that is coupled to the memory circuit 24and/or the control circuit 20 via the telemetry circuit 26 and asuitable communication link 27. Telemetry circuits used for this purposeare known in the art.

Alternatively, various other therapy circuits 23 may be utilized withthe cardiac detection system 21 of the present invention, such as othertypes of cardiac pacers or stimulators, implantable electricaldefibrillators, implantable monitoring devices, and the like.

In FIG. 5, a basic block diagram of the cardiac detection system 21 isshown. As seen in FIG. 5, the heart 10 and electrogram lead 12 arecoupled to signal conditioning circuitry 34. The signal conditioningcircuitry 34, in turn, is coupled to a morphology detector 36 and athreshold detector 38. The two detectors 36 and 38 are coupled to amicrocontroller 30. The microcontroller 30 is further coupled to thesignal conditioning circuitry 34 and to the therapy circuit 23.

The electrogram signal is transmitted to the signal conditioningcircuitry 34 via the electrogram lead 12 and is processed by the signalconditioning circuitry 34 into a plurality of discrete values. Thediscrete values are transmitted to the threshold detector 38 and to themorphology detector 36. The threshold detector 38 compares the amplitudeof each of the discrete values with an threshold value that is set bythe microcontroller 30. In the event that the amplitude of one of thediscrete values exceeds the event threshold value, an event sensedsignal is generated by the threshold detector 38. The event sensedsignal is then coupled to the microcontroller 30.

Simultaneously, the morphology detector 36 evaluates each of thediscrete values to determine the morphology of the electrogram signal.In response to this evaluation, the morphology detector 36 generates atleast one morphology value that is coupled to the microcontroller 30.The morphology value is used by the microcontroller 30 to adjust thesensitivity of the signal conditioning circuitry 34, e.g., either byadjusting the gain of an amplifier, the threshold or by adjusting thenarrow band filter bandwidth. Additionally, the morphology value is usedby the microcontroller 30 to generate an output signal. The outputsignal is coupled to the therapy circuit 23 over signal line 39. Suchoutput signal adjusts the therapy circuit 23 to deliver a more optimumtherapy to the heart 10 via the electrogram lead 12.

In FIG. 6, a detailed block diagram of the cardiac detection system 21is shown. The signal conditioning circuitry 34, threshold detector 38,morphology detector 36, microcontroller 30 and therapy circuit 23 areshown. The signal conditioning circuitry 34 is comprised of apre-amplifier 42 that is coupled to the electrogram lead 12; a narrowband filter 44 that is coupled to the pre-amplifier 42; and a 5-bit A/Dconverter 46, coupled to the output of the narrow band filter 44, whichconverts the filtered signal to a digital signal. The output of the A/Dconverter 46 is directed to the morphology detector 36 and the thresholddetector 38.

In addition, a wide band filter 60 and an 8-bit analog-to-digital (A/D)converter 61 are shown coupled between the signal conditioning circuitry34 and the microcontroller 30. Such wide band filter 60 and A/Dconverter 61 advantageously provide a signal path through which themicrocontroller 30 may directly monitor a digitized version of theelectrogram signal in addition to the processing of the electrogramsignal by the threshold detector 38 and the morphology detector 36.

The pre-amplifier 42 amplifies the electrogram signal from theelectrogram lead 12 using a gain amplitude that is controlled by a 3-bitgain set signal from the microcontroller 30. The gain set signal isgenerated in response to a control program discussed more completelybelow in reference to FIGS. 9-13.

Included in the electrogram signal are various noise signals, e.g.,undesirable components such as a T-wave, electromagnetic interference,myopotential voltage signals, and "baseline wander." (Baseline wander isa deviation, or drift, of the voltage of the electrogram signal duringwhat should be the electrically quiet part of the electrogram signal.)Each of these noise components have been recognized in the art. Afterthe electrogram signal is amplified, it is coupled to the narrow bandfilter 44. The narrow band filter 44 filters out most of the undesirablenoise components in the amplified electrogram signal, and removes theneed to compensate such signal for baseline wander in the controlprogram.

Sometimes it is desirable to widen or narrow the pass band of the narrowband filter 44, e.g., to adapt to different slew rates in theelectrogram signal that occur at the onset of an arrhythmia, as well asto compensate for variations due to different electrogram lead types andpositions. Advantageously, the pass bandwidth of the narrow band filter44 can be adjusted in response to a high pass set signal obtained fromthe microcontroller 30. The high pass set signal is generated by thecontrol program, or alternatively may be programmed to a desired valuevia the telemetry circuits and external programmer described above. Itis known that arrhythmia signals have lower frequency content, so whenthe signal is not a positively stable normal sinus rhythm, the bandwidthof the filter is widened.

The output of the narrow band filter 44 is coupled to the 5-bit A/Dconverter 46. In one embodiment, the 5-bit A/D converter 46 transformsthe amplified, filtered electrogram signal into one of thirty-twopossible 5-bit digital codes. The conversion rate of the converter isabout 1000 Hz, and the converter has two interleaved phases. During thefirst phase, a positive input produces a non-zero code (discrete value)indicative of the magnitude of the input, and a negative input producesa zero output. During the second phase, a positive input produces a zerooutput, and a negative input produces a non-zero output code (discretevalue) indicative of the magnitude of the input. The 5-bit A/D converter46 traverses one of these two phases during each conversion, andalternates between the first phase and the second phase duringsucceeding conversions.

The 5-bit converter 46 also provides a polarity signal indicating thepolarity of the last discrete value output therefrom, and anend-of-conversion signal indicating that an output code has been latchedon to a 5-bit output bus of the A/D converter 46. The 5-bit output bus,polarity signal and the end-of-conversion signal are coupled to thethreshold detector 38 and the morphology detector 36 over signal bus 37.

As seen in FIG. 6, the threshold detector 38 generates an event sensedsignal when an output from the 5-bit A/D converter 46 is larger than athreshold value. Note that this is an unsigned comparison: a discretevalue of either polarity (negative or positive) that exceeds thethreshold will cause the event sensed signal to be generated by thethreshold detector 38. The microcontroller 30 sets a starting thresholdvalue after the detection of a cardiac event, e.g., an R-wave. Thisstarting threshold value is a function of the previous detected peak andaverage values of the R-wave as described more fully below inconjunction with FIGS. 10 and 11. Additionally, the microcontroller 30prevents discrete values from being received by the threshold detector38 until a programmable (e.g.,0 to 50 ms) refractory time period haselapsed after the deliverance of the therapy to the heart 10 by thetherapy circuit 23. Such action prevents spurious re-detections of,e.g., a R-wave, after the therapy has been delivered to the heart 10.

One of the problems frequently encountered with some patients is a largeT-wave, representing the repolarization of the cardiac tissue. TheT-wave follows the depolarization of cardiac tissue, i.e., the R-wave.It is extremely important in a cardiac arrhythmia detection system thatthe T-wave not be confused with the R-wave. However, in patientsexhibiting a large T-wave, the T-wave may be of the same order ofmagnitude as the R-wave. Hence, the detection circuitry, if detection isbased solely on the electrogram signal exceeding a prescribed thresholdvalue, has no way to distinguish T-waves from R-waves.

The present invention advantageously addresses this problem by providinga threshold value to the threshold detector 38 that automaticallyadjusts the threshold value to a low value (increased sensitivity) inorder to best detect the R-wave, and (after the R-wave has beendetected) to a high value (decreased sensitivity) in order to best avoiddetection of the T-wave. After a sufficient time period at the decreasedsensitivity, selected to avoid detection of the T-wave, the thresholdvalue is then gradually decreased to its prior low value (i.e., thesensitivity is gradually increased) in anticipation of detecting thenext R-wave.

The manner of automatically adjusting the threshold value of thethreshold detector 38 in accordance with the present invention isgraphically depicted in FIG. 7. Represented in FIG. 7 are both anelectrogram signal 600, containing an R-wave 604 and a T-wave 606, and adynamic threshold value 602. Also included in FIG. 7 are "sample times",represented as tick marks along the horizontal axis. As describedpreviously, the electrogram signal 600 is sampled each sample time bythe A/D converter 46 (FIG. 6), with the resulting sampled value beingexamined by the threshold detector 38. If at any given sample time thevalue of the electrogram signal 600 is less than the dynamic thresholdvalue 602, then no cardiac event detection occurs. If, however, at agiven sample time the value of the electrogram signal 600 is greaterthan the dynamic threshold value 602, then a cardiac event detectionoccurs.

Thus, as shown in FIG. 7, at sample time 614, before the occurrence ofthe R-wave 604, the dynamic threshold value 602 is set to a low initialvalue (indicated as V_(i)) in anticipation of detecting an R-wave. Atsample time 614, the electrogram signal 600 is less than the dynamicthreshold value 602 so no R-wave detection has yet occurred. Betweensample times 616 and 618, the R-wave 604 begins, causing the electrogramsignal 600 to cross over the dynamic threshold value 602, therebysignaling the occurrence of the R-wave. As soon as the R-wave isdetected in this manner (i.e., at the next sample time 618 following thedetection of the R-wave), the dynamic threshold value 602 increases toits predetermined temporary value (indicated as V_(temp)) Suchpredetermined temporary threshold value is determined as a function ofthe detected average values (e.g., average peak, average max, etc.) ofprior cardiac events, as explained more fully below in conjunction withFIGS. 10 and 11. Alternatively, in some embodiments of the invention,the temporary threshold value and/or the initial threshold value may bea programmed value, selected by the physician at the time of implant orthereafter.

Once the dynamic threshold value 602 has been increased to thepredetermined temporary value, the dynamic threshold value 602 remainsfixed at such value for a predetermined, or programmable, period oftime, T_(F). Typically, T_(F) is defined to be a programmable number ofsample times of the A/D converter 46, e.g., 10 to 15 sample times (whichcorresponds to 10 to 15 ms, assuming a sample rate of 1000 Hz). Thetime, T_(F), during which the dynamic threshold value 602 remains fixedat the temporary threshold value is selected to be sufficiently long topass over a typical T-wave for the particular patient whose electrogramis being monitored.

At the conclusion of the time, T_(F), the dynamic threshold value 602ramps down, in a controlled manner, from the temporary value, V_(temp),to the initial value, V_(i). In a preferred embodiment, such rampingdown occurs by defining the dynamic threshold value 602 to be a numberof between 0 and 31, with 0 indicating the lowest (most sensitive)threshold value, and with 31 indicating the highest (least sensitive)threshold value. As the threshold value ramps down from its temporaryvalue to its initial value after the time period T_(F), it does so bydecrementing the dynamic threshold value 602 by one for a prescribednumber of sample periods. For example, as shown by the ramp down curve608 in FIG. 8, the dynamic threshold value is decremented by one foreach two sample times, until the dynamic threshold value reaches theinitial value.

In some instances, it is desirable to adjust the ramp down rate of thedynamic threshold value 602. For example, while the ramp down curve 608in FIG. 7 may be adequate for a given patient while at rest (having,e.g., a heart rate of 60 beats per minute, or one beat per second), itmay be totally inadequate if the patient's heart rate suddenly increasesbecause it may not be fully ramped down before the next R-wave occurs.Thus, the present invention includes a feature for automaticallyadjusting the ramp down rate as a function of the sensed heart rate. Ifthe sensed heart rate is slow, then a relatively slow ramp down rate canbe used, such as is represented by the ramp down curve 610 in FIG. 7(decrementing one unit for each four sample times). If the sensed heartrate is moderate, then a moderate ramp down rate can be used, such asthe ramp down curve 608 in FIG. 7 (decrementing one unit for each twosample times). If the sensed heart rate is fast, then a fast ramp downrate can be used, such as the ramp down curve 612 in FIG. 7(decrementing one unit for each sample time).

Advantageously, the adjustments of the threshold value, including theramp down rate, may occur under control of the microcontroller 30 (FIG.6) by simply providing the threshold detector 38 with a threshold setsignal. The threshold set signal includes, in a preferred embodiment, atleast three separate elements: a threshold set value; a thresholddecrement interval value; and a detect refractory set value. Thethreshold set value may comprise, e.g., 5 bits (corresponding to 32values). The threshold decrement interval value may comprise, e.g., 2bits (corresponding to four values). Such decrement interval valueeffectively sets the ramp down rate, as described above, therebyproviding four different ramp down rates. The detect refractory setvalue may comprise, e.g., 4 bits (corresponding to 16 values). Thedetect refractory set value effectively sets the duration of the timeT_(F) during which the threshold value remains fixed at the temporaryvalue. It is to be emphasized, of course, that the above values are onlyexemplary, and that any number of bits may be used to define theindicated variables in order to suit the needs of a particularapplication of the invention.

As evident to those of skill in the art, the actual implementation ofthe threshold detector 38 may occur in software (i.e., within themicrocontroller 30 and/or in hardware using, for example, appropriateregisters and logic gates). Such registers and logic gates may beconfigured in conventional manner to digitally compare two digitalnumbers, one of which is the sample value and the other of which is thethreshold value, and to decrement the threshold value as defined by thedetect refractory set value (fixed threshold time, T_(F)) and thethreshold decrement interval value (ramp down rate).

Graphically depicted in FIG. 14 is an alternate embodiment of thepreferred manner of automatically adjusting the threshold value of thethreshold detector 38. Represented in FIG. 14 are both an electrogramsignal 1600, containing an R-wave 1604 and a T-wave 1606, and a dynamicthreshold value 1602. As described previously, the electrogram signal1600 is sampled each sample time by the A/D converter 46, with theresulting sampled value being examined by the threshold detector 38. Ifat any given sample time the value of the electrogram signal 1600 isless than the dynamic threshold value 1602, then no cardiac eventdetection occurs. If, however, at a given sample time the value of theelectrogram signal 1600 is greater than the dynamic threshold value1602, then a cardiac event detection occurs.

As shown in FIG. 14, at sample time 1614, before the occurrence of theR-wave 1604, the dynamic threshold value 1602 is set to a low initialvalue (indicated as V_(i)) in anticipation of detecting an R-wave. Atsample time 1614, the electrogram signal 1600 is less than the dynamicthreshold value 1602 so no R-wave detection has yet occurred. Once theR-wave 1604 exceeds the dynamic threshold value 1602, the thresholddetector 38 indicates to the microcontroller 30 that an R-wave has begunand a delay period, T_(d), 1620 is initiated. The delay period, T_(d),1620 may be determined by the microcontroller 30, or other timingcircuitry, and may correspond to the programmed refractory period.

As soon as the R-wave is detected, the dynamic threshold value 1602 mayadjusted to a first temporary value (indicated as V_(temp1)) and a delayperiod, T_(d), 1620 is initiated. While the first temporary thresholdvalue is shown in FIG. 14 as being zero, it is preferably infinite sothat no detections occur. This may be accomplished by effectivelymasking the output of the threshold detector 38.

During the delay period, T_(d), 1620, the A/D converter 61 continues tosample the R-wave 1604 to determine the peak value. At the conclusion ofthe delay period, T_(d), 1620, the dynamic threshold value adjusts to asecond temporary value, V_(temp2), and begins to ramp down to theinitial value, V_(i). As the threshold value ramps, it does so bydecrementing the dynamic threshold value 1602 by one step (e.g., 200 μV)for a prescribed number of sample periods. For example, as shown by theramp down curve 1608 in FIG. 14, the dynamic threshold value isdecremented once for each sample time, until the dynamic threshold valuereaches the initial value.

In some instances, it is desirable to adjust the initial thresholdvalue, the ramp down rate, the first and second temporary values, andthe delay period, T_(d), of the dynamic threshold value 1602.Advantageously, the adjustments of the threshold value may occur undercontrol of the microcontroller 30.

In the preferred embodiment, the initial threshold value may be selectedfrom a percentage of the peak R-wave, or a fixed value, in the range of0.2 mV to 2.0 mV. The first temporary value is infinite, that is, thethreshold detector 38 is masked so that nothing is detected. The secondtemporary value is selectable in the range of 25% to 100% of the peakR-wave. The slope is also selectable in the range of 200 μV/10 ms to 200μV/125 ms and the delay period, T_(d), is selectable in the range of90-190 ms.

In a semi-automatic embodiment, the preferred values include: a delayperiod of 125 ms, a second threshold value of 50% of the peak R-waveamplitude, a ramp of 200 μV/62 ms and programmable selection of theinitial threshold value from 0.2-1.2 mV.

In a fully automatic embodiment, the preferred values include: a delayperiod of 125 ms, a second threshold value of 50% of the peak R-waveamplitude, a ramp of 200 μV/125 ms and an initial threshold value of 0.4mV.

It is to be emphasized, of course, that the above values are onlyexemplary, and that any number of bits may be used to define theindicated variables in order to suit the needs of a particularapplication of the invention.

With reference to FIG. 6, it is seen that the morphology detector 36 iscomprised of a plurality of detectors: a maximum detector 48, a minimumdetector 50, an accumulator (or summer) 52, a baseline sampler 54, and acounter (or baseline timer) 56. Each of these detectors is coupled tothe output bus 37 connected to the A/D converter 46.

The maximum detector 48 latches the value on the output bus 37 if: (1)its magnitude exceeds the value currently stored by the maximum detector48 (initially zero), and (2) the polarity output is positive, therebyalways trapping the most positive discrete value. The minimum detector50 latches the value on the output bus 37 if: (1) its magnitude exceedsthe value currently stored by the minimum detector 50 (initially zero),and (2) the polarity output is negative, thereby always trapping themost negative discrete value. Thus, together the maximum and minimumdetectors 48 and 50 completely bracket the excursion of the digitizedelectrogram signal. The maximum and minimum detectors 48 and 50 resetthemselves (i.e., reset the value currently stored in the detectors 48and 50 to zero) each time the microcontroller 30 reads them. This isnormally at the end of a cardiac cycle.

The accumulator 52 (labeled in FIG. 6 as an "unsigned summer") sums, ortotals, all of the discrete values produced by the A/D converter 46during the current cardiac cycle (i.e., since the previous event sensedsignal was generated by the threshold detector 38 without regard topolarity). Each time the A/D converter 46 issues an end-of-conversionsignal, the conversion (discrete value) is latched and added to aprevious accumulated magnitude value in the accumulator 52, therebyforming a new accumulated magnitude value. Thus, the accumulatedmagnitude value is representative of the total magnitude of all of thediscrete values issued by the A/D converter 46 during the currentcardiac cycle.

The baseline sampler 54 (labeled in FIG. 6 as an "unsigned baselinesampler") sums, or totals, all discrete values produced by the A/Dconverter 46 during a predetermined time period. The purpose of thebaseline sampler 54 is to sample what should be the electrically quiet(baseline) part of the electrogram signal that lies between intracardiacwaveform complexes. During normal sinus cardiac rhythm, the sumcollected by the baseline sampler 54 will be small and a sudden increasein the sum indicates the probability that a cardiac arrhythmia is inprogress.

The baseline sampler 54 (FIG. 6) is enabled by the baseline timer 56.The baseline timer 56 provides an enable signal that enables thebaseline sampler during the predetermined period of time. In oneembodiment, the predetermined period of time is defined by two set timesignals generated by the microcontroller 30: a set delay time signal T₁,and a set sample time signal, T₂. The set delay time signal determines adelay period after the detection of an event by the threshold detector38 during which the enable signal is not present and, hence, duringwhich the baseline sampler circuit is not enabled. The set sample timesignal defines a sample period or sample window after the delay periodduring which the enable signal is present and hence during which thebaseline sampler circuit 54 is monitoring the electrogram signal. Afterthe sample period or window, the enable signal is not generated. Byselecting appropriate values for the delay period, typically 100 to 150ms, and for the sample period or window, typically 20 to 50 ms, thebaseline sampler 54 is enabled during the quiet (baseline) part of theelectrogram signal. In an alternative embodiment, the average baselineis computed over an entire cycle length, since such calculation closelyapproximates the true, or quiescent, baseline.

While the baseline sampler 54 is enabled, it sums, or totals, alldiscrete values produced by the A/D converter 46. Each time the 5-bitA/D converter 46 issues an end-of-conversion signal, the conversion islatched and added to the previous total baseline value in the baselinesampler 54, thereby forming a new total baseline value.

A time-out detector, referred to as a cycle length counter 58 in FIG. 6,is also coupled to the microcontroller 30. The cycle length counter 58serves two functions. First, the cycle length counter 58 counts thenumber of end-of-conversion signals that are produced by the A/Dconverter 46 during a cardiac cycle, and thereby generates a cyclelength value. (Note, each end-of-conversion signal occurs at a knownrate (e.g., 1000 HZ) and thus has a known time period (e.g., 1 ms)associated therewith. Thus, the occurrence of 772 end-of-conversationsignals, for example, indicates a time period of 772 ms. Each time themicrocontroller 30 reads the cycle length counter 58, the cycle lengthcounter 58 is cleared (i.e., the timer begins to count from zero).Second, the cycle length counter 58 generates a cycle time-out signal ifthe count exceeds a preset number. The preset number (which correspondsto a preset time period) is set by the microcontroller 30 in accordancewith the control program or in accordance with a value stored in thememory circuit 24 via the telemetry circuit 26 as described above (FIG.4). The occurrence of the cycle time-out signal forces themicrocontroller 30 to read the morphology detector 36 in the event nocardiac event is sensed during the preset time period.

In FIG. 8, an exemplary intracardiac electrogram signal 450 is shown toillustrate a typical input signal to the cardiac detection system 21 andto further graphically represent a digitized electrogram signal 451generated in response to the electrogram signal 450 by the signalconditioning circuitry 34.

Referring first to the electrogram signal 450, it can be seen that theelectrogram signal is made up of five basic wave structures. First inthe electrogram signal 450 is the Q-wave, immediately followed by theR-wave, and then the S-wave. These three waves are commonly referred toas the QRS complex. The QRS complex is followed by the T-wave, and thenthe barely discernible P-wave. After passing through the narrow bandfilter 44 (FIG. 6), the T-wave and the P-wave are usually notdiscernible. After the P-wave, the cycle repeats beginning with the QRScomplex. The period, T₃, of the electrogram signal, from QRS complex toQRS complex, is referred to as the cardiac cycle. The period T₃ for theelectrogram signal 450 shown in FIG. 8 may be in the range ofapproximately 800 to 300 ms (corresponding to a rate of approximately 70to 200 ppm).

The digitized electrogram signal 451, also shown in FIG. 8, reveals thatthe 5-bit A/D converter produces a non-zero output code 414 during afirst sample period, approx. 1 ms, because the input electrogram signal450 is positive when the electrogram signal is digitized and the A/Dconverter 46 is in the first phase, as mentioned above. During a secondsample period, an output code 420 is produced that is zero because theinput electrogram signal is still positive and the A/D converter 46 isin the second phase, as mentioned above. This process continues withevery other sample period producing a non-zero output code (indicativeof a positive input electrogram signal) until the occurrence of theQ-wave. With the occurrence of the Q-wave, a negative excursion of theelectrogram signal first occurs, as indicated at 404. Immediatelyfollowing the negative excursion 404, a positive excursion occurs at401, representative of the positive R-wave being detected. The amplitudeof the digitized electrogram signal at 401 exceeds the threshold valuerepresented by line E--E. Exceeding the threshold E--E causes thethreshold detector 38 to generate the event sensed signal describedabove.

When the R-wave is detected, a delay period, T₁, begins during which thebaseline sampler 54 is disabled. Following the delay period, T₁, asample period, T₂, begins during which the baseline sampler 54 isenabled. As explained above, the baseline sampler is disabled after thesample period, T₂, until the next sample period begins. The next delayperiod begins after the next event sensed signal is generated at 402.This will occur when the digitized electrogram signal again exceeds thethreshold value designated by line E--E.

It should be noted that the sample periods shown in FIG. 8 are, forreasons of clarity, not drawn to scale relative to the horizontal (time)axis. A typical cardiac cycle, for example, may be 1000 ms(corresponding to 60 heartbeats per minute). At a sample rate of 1000Hz, one sample would be made every 1 ms, or 1000 samples would be madeduring a cardiac cycle. For purposes of clarity, however, only about 48to 50 samples are shown in FIG. 8 as being taken during the cardiaccycle, T₃.

In FIGS. 9, 10, 11, 12 and 13, a flowchart of a control program used bythe microcontroller 30 is shown. Each main step of the flowchart isshown as a "box" or "block," with each box or block having a referencenumeral associated therewith. The control program is called by themicrocontroller 30 whenever the microcontroller 30 receives the eventsensed signal from the threshold detector 38 or receives the cycletime-out signal from the cycle length detector 58 (Block 1001). Underthe direction of the control program, the microcontroller 30 reads theminimum value, the maximum value, the total baseline value and the cyclelength value (Block 1003) of the electrogram signal during a givencardiac cycle. Note that the maximum and minimum detectors, the baselinesampler, and the cycle length detector are reset (to zero) after theyare read by the microcontroller 30. The average baseline value is thencalculated by dividing the total baseline value by the number of sampledelay values (Block 1005). Next the average baseline value is subtractedfrom the maximum value (Block 1007) and from the minimum value (Block1009), thereby removing any normally distributed noise level. A variablecalled current polarity is set equal to the polarity value of thediscrete value that caused the event sensed signal to be asserted.

The following discussion assumes that the current polarity variable ispositive, however, symmetrical operation would occur if the currentpolarity variable were negative. The current polarity is tested (Block1011) and, if the polarity is positive, execution proceeds at Block 1013(FIG. 10). The magnitude of the maximum value is compared with theminimum value (Block 1015). During normal sinus rhythm, the maximumvalue will be larger than the minimum value whenever the currentpolarity variable is positive. If the maximum value is smaller than theminimum value, a radical change in the morphology of the electrogramsignal is indicated. In response to this radical change, the MORPHhd--CHANGE variable is set to "POLARITY" (Block 1029), the polarityvariable is set to negative (Block 1031), the variable MIN₋₋ AVERAGE isset equal to the current minimum value (Block 1033) and the variable MAXAVERAGE is set equal to the current maximum value (Block 1035). It isimportant that the MIN₋₋ AVERAGE and the MAX₋₋ AVERAGE variables be setto the current minimum and maximum values, respectively, because thereis no reason to assume that these average variables are meaningful aftersuch a radical change in morphology. The generation of MIN₋₋ AVERAGE andMAX₋₋ AVERAGE is discussed below. Execution of the program continues inBlock 1037 (FIG. 11).

If the maximum value is larger than the minimum value (as determined atBlock 1015 of FIG. 10), the maximum value is compared to the MAX₋₋AVERAGE variable (Blocks 1017 and 1019). If the maximum value is smallerthan one-half times the MAX₋₋ AVERAGE variable, the variable MORPH₋₋CHANGE is set to "DECREASE" (Block 1025); and if the maximum value isgreater than one and one-half times the MAX₋₋ AVERAGE variable, thevariable MORPH₋₋ CHANGE is set to "INCREASE" (Block 1027). If themaximum value is between one-half times the MAX₋₋ AVERAGE, and one andone-half times the MAX₋₋ AVERAGE, the variable MORPH₋₋ CHANGE is set to"NONE" (Block 1021).

Next, the variable MAX₋₋ AVERAGE (Block 1023) is updated using thefollowing relationship:

MAX₋₋ AVERAGE=3/4MAX₋₋ AVERAGE)×1/4(maximum value)

The threshold value (Block 1095) is generated according to the followingrelationship:

THRESHOLD VALUE=1/2(MAX₋₋ AVERAGE)+(average baseline value)

Note that the variable MIN₋₋ AVERAGE is similarly generated and used togenerate the threshold value if the current polarity variable isnegative (FIG. 11). Execution continues at Block 1097 (FIG. 12). The setTHRESHOLD is then used by the threshold detector 38 as describedpreviously. That is, the THRESHOLD value is written to the thresholddetector 38 as a threshold set signal (Block 1063, FIG. 9), andexecution of the control program terminates (Block 1065) until anotherevent sensed signal is coupled to the microcontroller or until a cycletime-out signal is received.

Prior to writing the THRESHOLD value to the event detector, anarrhythmia flag value is generated by dividing the maximum value (orpeak value) by the average baseline value (Block 1101) (FIG. 12). Thearrhythmia flag value may be used, e.g., by a defibrillation program toset an arrhythmia flag signal, or otherwise determine whether or nothigh energy shock therapy should be delivered to the heart. Next, thevariable MAX₋₋ AVERAGE is tested. If MAX₋₋ AVERAGE is greater thantwenty-four (Block 1103) then the gain of the amplifier is decreased byone-half (Block 1109) and the variable GAIN₋₋ CHANGE is set to"DECREASE" (Block 1115). If MAX₋₋ AVERAGE is less than eight (Block1105), then the gain of the amplifier is doubled (Block 1111) and thevariable GAIN₋₋ CHANGE is set to "INCREASE" (Block 1113). If the MAX₋₋AVERAGE is between twenty-four and eight, the variable GAIN₋₋ CHANGE isset to "NONE" (Block 1051). (NOTE: For a 5 bit A/D converter which has apossible 0-31 conversion codes, clipping would occur at conversion code31 and no signal correlates to conversion code 0. Thus, conversion code8 and 24 are chosen in Blocks 1103, 1105 as arbitrary limits todetermine if the signal is within an optimum dynamic range.) Executionof the control program then continues at Block 1061 (FIG. 9). The GAIN₋₋CHANGE variable is used by the pre-amplifier 42 to set its gain, asdescribed previously.

As mentioned above, similar operation of the control program isexhibited in the event that the current polarity is negative. Theprimary difference in the execution of the control program when thecurrent polarity is negative is that the minimum value is used togenerate the variable MORPH₋₋ CHANGE and the threshold value (Blocks1069 through 1091 of FIG. 11); and to generate the variable GAIN₋₋CHANGE, and set the gain of the amplifier (Blocks 1045 through 1061 ofFIG. 13).

The table, below, summarizes the signals used to indicate a change fromrhythmic to arrhythmic cardiac activity, and vice verse. For example, ifthe peak R-wave signal divided by the average baseline is less than aprescribed threshold, then normal sinus rhythm is present. However, onceit exceeds a prescribed threshold, then there is a high confidence levelthat an arrhythmia is present. Similarly, if the morphology changevalue, POLARITY, is set (indicating that the average amplitude of anR-wave has reversed polarity), or if there is an INCREASE or DECREASE inamplitude, or a GAIN change, then there is a high confidence level thatR-waves are now being generated from a new, ectopic, location. Thus, thepresent invention indicates to the microcontroller that an arrhythmia ispresent. Furthermore, combinations of these morphology change signalscan be used together, or in combination with conventional detectionmethods (e.g., tachycardia rate threshold, sudden onset, stability,etc.) to produce even higher confidence levels that an arrhythmia hasbegun.

    ______________________________________                                        AVG. BASELINE <threshold value                                                                             Normal rhythm                                    AVG. BASELINE >threshold value                                                                             arrhythmia                                                                    present                                          PEAK ÷ AVG.                                                                             <threshold value                                                                             Normal rhythm                                    BASELINE                                                                      PEAK ÷ AVG.                                                                             >threshold value                                                                             arrhythmia                                       BASELINE                     present                                          morphology change                                                                           POLARITY (changed)                                                                           arrhythmia                                       value                        present                                          morphology change                                                                           INCREASED      arrhythmia                                       value         amplitude      present                                          morphology change                                                                           DECREASED      arrhythmia                                       value         amplitude      present                                          morphology change                                                                           GAIN (changed) arrhythmia                                       value                        present                                          morphology change                                                                           NONE (no change in                                                                           Normal rhythm                                    value         amplitude, gain,                                                              or polarity)                                                    ______________________________________                                    

As described above, it is thus seen that the present invention providesa way of automatically adjusting the sensitivity (gain and/or threshold)of a cardiac event detection system in response to varying electrogramsignals.

It is also seen that the invention provides implantable circuitry thatinexpensively acquires information from an electrogram signal over apotentially long (e.g., infinite) period of time, without having to useexternal storage devices, e.g., magnetic tape recorders.

It is further seen that the invention provides an implanted system thatadvantageously acquires information from an electrogram signal withoutthe need for large memories to store all or portions of the electrogramsignal.

Additionally, it is seen that the invention provides an implantablearrhythmia detection system that generates output signals useable by animplantable therapy circuit, such as a pacemaker, cardioverter, ordefibrillator, or the like, and that facilitates automatic control andadjustment of the therapy provided by such therapy circuit.Advantageously, it is seen that the output signals used to control thetherapy circuit are readily generated without the need for performingcomplicated and highly power-consuming mathematical computations in animplanted microcontroller.

While the invention herein disclosed has been described by means ofspecific embodiments and applications thereof, numerous modificationsand variations could be made thereto by those skilled in the art withoutdeparting from the scope of the invention set forth in the claims.

What is claimed is:
 1. An implantable cardiac event detection system foreliminating sensing of high amplitude T-waves, the system including animplantable lead that transmits an electrogram signal from a heart,comprising:an amplifier, coupled to the implantable lead, for amplifyingan electrogram signal; threshold detection means, coupled to theamplifier, for detecting an occurrence of an R-wave in the electrogramsignal whenever the magnitude of the electrogram signal exceeds athreshold value; timing means for generating a delay period followingthe detection of an R-wave; and control means, responsive to thedetection of the R-wave, for dynamically changing the threshold valuefrom an initial value to a temporary value following the delay period,and then gradually returning the threshold value from the temporaryvalue back to the initial value in accordance with a prescribed slope,so that a high amplitude T-wave that follows the R-wave is not sensed.2. The implantable cardiac event detection system, as set forth in claim1, wherein the control means changes the threshold value from thetemporary value back to the initial value in a predetermined stepwisefashion.
 3. The implantable cardiac event detection system, as set forthin claim 2, wherein the prescribed slope is selectable from the range ofslopes consisting of 200 μV/10 ms to 200 μV/125 ms.
 4. The implantablecardiac event detection system, as set forth in claim 1, furthercomprising:means for programmably adjusting the delay period beforechanging the threshold value from its temporary value back to itsinitial value.
 5. The implantable cardiac event detection system, as setforth in claim 1, wherein the delay period is programmable between 90and 190 ms.
 6. The implantable cardiac event detection system, as setforth in claim 5, wherein the delay period is approximately 125 ms. 7.The implantable cardiac event detection system, as set forth in claim 1,further comprising:means for programmably selecting at least one of theinitial threshold value and the temporary threshold value.
 8. Theimplantable cardiac event detection system, as set forth in claim 7,wherein the percentage of the peak R-wave amplitude value is 50%.
 9. Theimplantable cardiac event detection system, as set forth in claim 8,further comprising:means for operating in an automatic mode, wherein theprescribed slope is 200 μV/125 ms and the initial threshold voltage isapproximately 0.4 mV.
 10. The implantable cardiac event detectionsystem, as set forth in claim 8, further comprising:means for operatingin a semi-automatic mode, wherein the prescribed slope is 200 μV/62 msand the initial threshold voltage is programmable between 0.2 mV and 1.2mV.
 11. The implantable cardiac event detection system, as set forth inclaim 1, wherein the control means comprises:peak detecting means fordetecting a peak R-wave amplitude value; means for setting at least oneof the initial threshold value and the temporary value to a percentagein the range of 25% to 100% of the peak R-wave amplitude value.
 12. Animplantable cardiac event detection system having an improveddual-sensing means for detecting cardiac arrhythmias, the systemincluding an implantable lead that transmits an electrogram signal froma heart, the improvement comprising:an amplifier, coupled to theimplantable lead, for amplifying an electrogram signal; thresholddetection means, coupled to the amplifier, for detecting an occurrenceof an R-wave in the electrogram signal whenever the magnitude of theelectrogram signal exceeds a threshold value; control means forprogrammably selecting one of a semi-automatic sensing mode and anautomatic sensing mode for the detection of cardiac arrhythmias, and forprogrammably changing the threshold value from an initial value to atemporary value in response to a detected R-wave, the initial valuebeing programmable in the semi-automatic mode and fixed in the automaticmode, and then gradually returning the threshold value from thetemporary value back to the initial value in accordance with a desiredslope as defined by the programmed sensing mode, so that a highamplitude T-wave that follows the R-wave is not sensed and arrhythmiascan be reliably detected in one of the automatic and the semi-automaticmode.
 13. The improvement, as set forth in claim 12, furthercomprising:timing means for generating a delay period following thedetection of an R-wave; and means for triggering the thresholdadjustment means to change the threshold value from the initial value tothe temporary value after the delay period has terminated.
 14. Theimprovement, as set forth in claim 13, wherein the delay period isprogrammable between 90 and 190 ms.
 15. The improvement, as set forth inclaim 12, wherein the desired slope is in the range of 200 μV/10 ms to200 μV/125 ms.
 16. The improvement, as set forth in claim 12, furthercomprising:a peak detector, coupled to the implantable lead, fordetecting the peak R-wave amplitude of the electrogram signal; and meansfor setting the temporary value to a percentage in the range of 25% to100% of the peak R-wave amplitude value.
 17. The improvement, as setforth in claim 12, further comprising:a peak detector, coupled to theimplantable lead, for detecting the peak R-wave amplitude of theelectrogram signal; and means for setting the temporary value to atleast 50% of the peak R-wave amplitude value.
 18. The improvement, asset forth in claim 17, further comprising:timing means for generatingdelay period following the detection of an R-wave, the delay periodbeing approximately 125 ms in duration; and means for triggering thethreshold adjustment means to change the threshold value from theinitial value to the temporary value after the delay period hasterminated.
 19. The improvement, as set forth in claim 18, wherein theprogrammable initial value in the semi-automatic mode is programmable inthe range of 0.2 mV and 1.2 mV.
 20. The improvement, as set forth inclaim 19, wherein the desired slope in the semi-automatic mode isapproximately 200 μV/62 ms.
 21. The improvement, as set forth in claim19, wherein the desired slope in the automatic mode is approximately 200μV/125 ms.
 22. The improvement, as set forth in claim 18, wherein thefixed initial value in the automatic mode is approximately 0.4 mV.